1. Field of the Invention
This invention relates broadly to medical electrical leads for electrical stimulation or electrical sensing of body organs or tissues and their method of fabrication. More particularly, this invention relates to implantable cardiac leads for delivering electrical stimulation to the heart, e.g., pacing pulses and cardioversion/defibrillation shocks, and/or sensing the cardiac electrogram (EGM) or other physiologic data.
2. State of the Art
Implantable medical electrical stimulation and/or sensing leads (referred to herein as “pacemaker leads or lead(s)”) are well known in the fields of cardiac stimulation and monitoring, including cardiac pacing and cardioversion/defibrillation. In these applications, a pacemaker or cardioverter/defibrillator implantable pulse generator (IPG) or a cardiac monitor is coupled to the heart through one or more of such leads. The proximal end of such leads is formed with a connector element which connects to a terminal of the IPG or cardiac monitor. The distal end of such leads includes a distal stimulation and/or sensing electrode that is fixated to tissue at the desired treatment site. A lead body extends between the distal and proximal ends. The lead body comprises one or more electrical conductors surrounded by an insulating outer sleeve. Each electrical conductor provides an electrical signal path between the proximal connector element (and the IPG or cardiac monitor coupled thereto) and the distal stimulation and/or sensing electrode. A lead having a single distal stimulation and/or sensing electrode is typically referred to as a unipolar lead. A lead having two or more distal stimulation and/or sensing electrodes is typically referred to as a bipolar (or a multi-polar) lead. The leads are typically implanted using an endocardial approach or an epicardial approach. The endocardial approach is the most common method. The epicardial approach is a less common method in adults, but more common in children.
In the endocardial approach, a local anesthetic is typically applied to numb an incision area of the chest (typically adjacent the collar bone) where one or more leads and the IPG or cardiac monitor are inserted. Each lead is inserted through the incision and into a vein, then guided through a transvenous pathway to the heart with the aid of fluoroscopy. The distal lead electrode is affixed to the heart muscle at the desired treatment site. The proximal connector element of the lead is coupled to the IPG or cardiac monitor, and the IPG or cardiac monitor is placed in a pocket created under the skin in the upper chest. The transvenous pathway can include a number of twists and turns, and the lead body can be forced against bony structures of the body that apply stress to it.
The epicardial approach requires open heart surgery wherein the distal lead electrode of one or more leads is affixed directly to the heart tissue at the desired treatment site, instead of inserting the lead(s) through a vein. The proximal connector element of the lead is coupled to the IPG or cardiac monitor, and the IPG or cardiac monitor is placed in a pocket created under the skin in the abdomen.
In all applications, the heart beats approximately 100,000 times per day or over 30 million times a year, and each beat stresses at least the distal portion of the lead body. The lead conductors and insulation are subjected to cumulative mechanical stresses, as well as material reactions as described below, that can result in degradation of the insulation or fractures of the lead conductors with untoward effects on device performance and patient well being.
In order to facilitate advancement through the transvenous pathway (for the endocardial approach) and minimize stress on the lead body (for all applications), flexible lead bodies have been developed using smaller diameter coiled wire conductors and flexible insulating materials, most notably polyurethane compositions. However, problems have been encountered as to the bio-stability of such lead materials. More particularly, it is acknowledged that there are a number of mechanisms for degradation of elastomeric polyurethane insulation of the lead body in vivo. One is environmental stress cracking (ESC), which is the generation of crazes or cracks in the polyurethane elastomer produced by the combined interaction of a medium capable of acting on the elastomer and a stress level above a specific threshold. Another is metal ion induced oxidation (MIO) in which polyurethane elastomers exhibit accelerated degradation from metal ions such as cobalt ions, chromium ions, molybdenium ions and the like which are used alone or in alloys in the conductive wire of the lead body.
The degradation mechanism of polyether urethanes was elucidated by Anderson's group at Case Western Reserve University (Cleveland, Ohio). They found that the carbon alpha to the ether of the polyether soft segment was oxidized to ester either by superoxide (O3) produced by polymorphonuclear leucocytes (PMNs) and the like, or by metal ion contact of the polyurethane, as occurs on the inside of pacemaker lead insulators. Subsequent hydrolysis of the ester cleaves the macromolecule, and in the presence of flexion, cracks develop. Realizing that the ether groups were vulnerable, the inventor of the subject application introduced more biostable polycarbonate urethanes for implant applications, which were initially commercialized under the trade name Corethane™ by Corvita Corp. of Miami, Fla. and now commercialized under the name Bionate® by DSM PTG of Berkley, Calif.
The improved biostability of polycarbonate urethanes was confirmed by Stoke's group at Medtronic using the “Stokes Test”, in which a tube of the material is stretched over a dumbbell-shaped mandrel and exposed to oxidizing and hydrolyzing chemicals, or is implanted in the body for a predetermined time. Materials that are readily susceptible to oxidation and hydrolysis crack in this model; significantly, the polycarbonate urethanes did not crack over the duration tested.
Although polycarbonate urethanes demonstrated superior biostability relative to polyether and polyester urethanes, they too eventually exhibited biodegradation as manifested by surface cracking The fractures were most noticeable in areas with large numbers of macrophages on histology. Importantly, Wilson's group (The Hospital for Sick Children, Toronto, Ontario) also observed that these degrading implants attracted a plethora of polymorphonuclear leukocytes, especially during the early weeks of implantation. Further, the cleaner the polycarbonate urethane (less extractables, washed surfaces), the more intense the inflammation. Further observations were the attraction of macrophages, foreign body giant cells and the phagocytosis of small “chunks” of polyurethane. Lastly, it was also observed upon careful examination that crack formation in microfilamentous grafts as early as 1 month after implantation. A summary of these finding was recorded in the article by Pinchuk et al. entitled “Medical applications of poly(styrene-block-isobutylene-block-styrene) (“SIBS”),” Biomaterials (2007), doi:10.1016/j.biomaterials.2007.09.041. In summary, polyurethanes exhibit degradation with time with signs of the problem occurring within weeks of implantation. Degradation is due to oxidation, most likely by superoxide produced by phagocytes (“scavenger cells”); the more degradation, the greater number of scavenger cells that migrate to the site, the worse the degradation. The more oxygen that can penetrate the polyurethane, the more it degrades and similarly, the more water that absorbs into the polyurethane, the better the transport of oxygen and other substances, for example, hydrogen ion, that can degrade the polymer.